When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0) applied along the z axis of a Cartesian coordinate system, the individual magnetic moments of the spins in the tissue attempt to align with this polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, Mz, may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic moment Mt. A NMR signal is emitted by the excited spins after the excitation signal B1 is terminated, this signal may be received and processed to form an image or produce a spectrum.
When utilizing these signals to produce images, magnetic field gradients (Gx, Gy and Gz) are employed. Typically, the region to be imaged is scanned by a sequence of measurement cycles in which these gradients vary according to the particular localization method being used. The resulting set of received NMR signals are digitized and processed to reconstruct the image using one of many well known reconstruction techniques.
Radio frequency antennas, or coils are used to produce the excitation field B1 and other RF magnetic fields in the subject being examined. Such coils are also used to receive the very weak NMR signals that are produced in the subject. Such coils may be so-called “whole body” coils that are large enough to produce a uniform magnetic field for a human subject or, they can be much smaller “local” coils that are designed for specific clinical applications such as head imaging, knee imaging, wrist imaging, etc. Local coils may be either volume coils or surface coils.
The most common whole body coil found in commercial MRI systems is the so-called “birdcage” coil first disclosed in U.S. Pat. Nos. 4,692,705; 4,694,255; and 4,680,548. A birdcage coil has a pair of circular end rings which are bridged by a plurality (typically 8 to 24) of equi-spaced longitudinal straight segments. In a primary mode, currents in the straight segments are sinusoidally distributed which results in good B1 field uniformity across the axis of the coil. Birdcage coils are the structure of choice in horizontal field MRI systems because they produce a homogeneous magnetic B1 field in the bore of the magnet. When properly designed and constructed, they have a high SNR which enables them to pick up the small NMR signals emanating from the subject under examination.
The birdcage coil is tuned by proper selection of capacitors which are distributed along the lengths of the straight segments, distributed around each end ring or both. Matching and tuning are commonly achieved by connecting variable capacitors in an “L” configuration at the drive ports. Birdcage coils are typically driven at one, two, or more recently, four ports. Multi-port drive, where each drive source is appropriately phased, ensures uniform, circularly polarized B1 fields in the imaging volume at B0 field strengths of 1.5 T or less. Efforts to improve the tunability of birdcage coils either provide fewer capacitor adjustments that distort the homogeneity of the B1 field or provide expensive and complex tuning structures such as those described in U.S. Pat. Nos. 6,396,271 and 6,236,206.
High field MRI, with B0 field strength of 3 T or higher, is rapidly winning acceptance in both clinical and research programs. High field MRI offers many benefits, while simultaneously presenting many research and design problems. The main benefit of high field MRI is increased signal to noise ratio (SNR). SNR increases linearly with static field strength. This increase provides significant advantages in terms of spatial, temporal, and spectral resolution. Functional MRI (fMRI) is an application of MRI that is used to analyze brain function using blood oxygenation level dependent (BOLD) contrast to detect the brain's response to a specific stimulus. For accurate analysis, fMRI demands both high spatial and high temporal imaging resolution. Several studies have demonstrated the benefit of high static field strength for fMRI applications. Proton MR Spectroscopy (MRS) is used for metabolic characterization of tumors, and for monitoring treatment of epilepsy, stroke, infections, and multiple sclerosis, though this list is by no means exhaustive. Since a higher static magnetic field provides a greater absolute chemical shift, the resolution of metabolite peaks improves, leading to greater accuracy in the identification and quantification of metabolites. These benefits also apply to imaging and spectroscopy with carbon-13, whether thermal or hyperpolarized. MR microscopy is used for imaging at sub-millimeter resolution. This technique also benefits from a higher B0.
The main problem with high field imaging is B1 (RF) field inhomogeneity induced by the sample under examination. The dielectric properties of the sample are largely responsible for these effects. This phenomenon may be seen in a high dielectric phantom imaged at high field; a characteristic bright spot is seen, surrounded by dark bands. Human tissues generally have a relative permittivity ranging between 50 and 80 at 128 MHz (3 T field, proton frequency). This is compounded by the fact that the human body is inhomogeneous, unlike a typical MRI phantom. Thus, the wavelength of RF in the body at 128 MHz ranges from around 26 cm to 34 cm. These wavelengths are comparable to the dimensions of the human body; hence phase shifts occur in the body. Due to the multiple sources of B1 field in a coil (multiple conductors), these phase shifts result in interference patterns inside the body. These patterns manifest themselves on MR images. As one increases the static magnetic field above and beyond 3 T, this problem worsens. One solution to this problem is to “pre-distort” the RF B1 field in such a way as to compensate for the effects of an inhomogeneous dielectric. It is therefore desirable to design RF coils with provision for phase and amplitude control for each conductive element.
The sample under examination has conductive properties along with the dielectric properties mentioned above. As a result, stray or “parasitic” capacitance is created between the load and the conductive elements of the RF coil. For coils designed to operate at static fields of 1.5 T and below, the lumped capacitive elements used to resonate the coil at the desired Larmor frequency are fairly large compared to the stray capacitances introduced by the specimen or load. At higher B0 fields, parasitic capacitances become comparable in size to coil component capacitors. It follows that an asymmetric load placed in the coil would lead to unequal frequency shifts in the resonant loops of the coil, thereby perturbing the ideal sinusoidal current distribution and eliminating the desired “neutral point” in the center of the coil, leading to poor B1 field homogeneity. In addition, local E field “hot spots” may occur, which may cause excessive RF heating above the specific limits on specific absorption rates (SAR) for human tissue set by the FDA. This problem can be ameliorated somewhat by using coils designed around transmission line elements, which have capacitance distributed along their length, in contrast with the lumped capacitance elements used in the birdcage coil designs.
Examples of distributed capacitance coils include the original transverse electromagnetic (TEM) resonator described by Roschmann in U.S. Pat. No. 4,746,866, and the TEM volume resonator described by Vaughan in U.S. Pat. No. 5,557,247. In U.S. Pat. No. 5,557,247, an array of transmission line elements with adjustable, re-entrant center conductor elements are arranged on a circular perimeter so as to form a multi-mode resonant cavity structure. In practice, the device is tuned and matched such that a specific mode coincides with the Larmor frequency. Circular polarized versions of this coil are driven at two ports 90 degrees out of phase with one another, with the remaining conductive elements couple to one another by induction. The phase increment between adjacent elements is N/360 degrees, where N is the number of elements. For proper operation of the coil, all elements must be of the same impedance. In practice, it is often necessary to adjust the re-entrant center conductors of each element in order to achieve this condition.
There are a number of clinical applications where MR images are acquired at different Larmor frequencies. Hydrogen (H1) is the spin species of choice for most MR imaging applications, but other paramagnetic spin species such as phosphorus (31P), fluorine (19F), carbon (13C), sodium (23Na), helium (3He) and xenon (129Xe) are also employed. Most of these alternative spin species are of interest in MR spectroscopy, but the use of helium for imaging the lung, for example, has significant clinical potential. As indicated above, the birdcage coil is difficult to tune at more than one Larmor frequency and the substantial change in Larmor frequency required to examine these alternative spin species is not practical.
Multinuclear excitation and reception coils have been proposed. In U.S. Pat. No. 4,799,016 for example, two birdcage coils are formed on one cylindrical substrate, with one coil tuned to hydrogen (1H) and the other tuned to phosphorus (13P). To reduce interaction between the coils, the fields they produce are offset 90° in phase. In U.S. Pat. No. 5,990,681 an RF coil is described which has an adjustment end ring provided on the end of a birdcage coil, wherein the ring can be rotated to change its Larmor frequency. An important limitation of prior multinuclear coils is that they consist of multi-modal resonant structures such as birdcage or TEM volume resonators. If one of the resonant modes corresponding to the Larmor frequency of the first nucleus coincides with the fundamental resonant mode corresponding to the Larmor frequency of the second nucleus, the isolation between the two components of the multi-nuclear coil degrades, and the two components of the coil cannot be operated simultaneously. In addition, poor isolation tends to degrade efficiency for each component of the coil in question. In practice, this means that when an image of a subject is acquired at the Larmor frequency of one nucleus, a subsequent scan must be performed if an image is to be obtained at the Larmor frequency of the second nucleus. During the time interval between scans, subject motion may occur, making the co-registration of the two scans difficult. It is therefore desirable to design multi-nuclear coils wherein the component coils are not multi-modal in nature, and the component coils have good electrical isolation and nearly identical spatial profiles.